Photoacoustic remote sensing (pars)

ABSTRACT

A photoacoustic remote sensing system (PARS) for imaging a subsurface structure in a sample has an excitation beam configured to generate ultrasonic signals in the sample at an excitation location; an interrogation beam incident on the sample at the excitation location, a portion of the interrogation beam returning from the sample that is indicative of the generated ultrasonic signals; an optical system that focuses at least one of the excitation beam and the interrogation beam with a focal point that is below the surface of the sample; and a detector that detects the returning portion of the interrogation beam.

CROSS-REFERENCE TO RELATED APPLICATION(S)

This patent application is a continuation of U.S. patent applicationSer. No. 16/116,038, filed on Aug. 29, 2018, which is a continuation ofU.S. patent application Ser. No. 14/919,565, filed on Oct. 21, 2015, nowU.S. Pat. No. 10,117,583, which claims priority to U.S. ProvisionalApplication No. 62/067,197, filed Oct. 22, 2014, the entireties of whichare incorporated herein by reference.

BACKGROUND Technical Field

This relates to the field of biomedical optics imaging and, inparticular, to a laser and ultrasound-based method and system for invivo or ex vivo, non-contact imaging of biological tissue.

Description of the Related Art

Photoacoustic imaging is an emerging hybrid imaging technology providingoptical contrast with high spatial resolution. Nanosecond or picosecondlaser pulses fired into tissue launch thermo-elastic-induced acousticwaves which are detected and reconstructed to form high-resolutionimages. Photoacoustic imaging has been developed into multipleembodiments, including photoacoustic tomography (PAT), photoacousticmicroscopy (PAM), optical-resolution photoacoustic microscopy (OR-PAM),and array-based PA imaging (array-PAI). In photoacoustic tomography(PAT) signals are collected from multiple transducer locations andreconstructed to form a tomographic image in a way similar to X-ray CT.In PAM, typically, a single element focused high-frequency ultrasoundtransducer is used to collect photoacoustic signals. A photoacousticsignal as a function of time (depth) is recorded for each position in amechanically scanned trajectory to form a 3-D photoacoustic image. Themaximum amplitude as a function of depth can be determined at each x-yscan position to form a maximum amplitude projection (MAP) C-scan image.Photoacoustic microscopy has shown significant potential for imagingvascular structures from macro-vessels all the way down tomicro-vessels. It has also shown great promise for functional andmolecular imaging, including imaging of nanoparticle contrast agents andimaging of gene expression. Multi-wavelength photoacoustic imaging hasbeen used for imaging of blood oxygen saturation, by using known oxy-and deoxy-hemoglobin molar extinction spectra.

In traditional photoacoustic imaging, spatial resolution is due toultrasonic focusing and can provide a depth-to-resolution ratio greaterthan 100. In OR-PAM, penetration depth is limited to ˜1 mm in tissue(due to fundamental limitations of light transport) but resolution ismicron-scale due to optical focusing. OR-PAM can provide micron-scaleimages of optical absorption in reflection-mode, in vivo, something thatno other technique can provide. OR-PAM is capable of imaging bloodvessels down to capillary size noninvasively. Capillaries are thesmallest vessels in the body and so much crucial biology occurs at thislevel, including oxygen and nutrient transport. Much can go wrong at thecapillary level too. In cancers, cells have an insatiable appetite foroxygen and nutrients to support their uncontrolled growth. They invoke arange of signaling pathways to spawn new vessels in a process known asangiogenesis and these vessels typically form abnormally. Tumors areoften highly heterogeneous and have regions of hypoxia. Photoacousticimaging has demonstrated the ability to image blood oxygen saturation(SO2) and tumor hypoxia in vivo.

In most photoacoustic and ultrasound imaging systems, piezoelectrictransducers have been employed, in which an ultrasound coupling mediumsuch as water or ultrasound gel is required. However for many clinicalapplications such as wound healing, burn diagnostics, surgery, and manyendoscopic procedures physical contact, coupling, or immersion isundesirable or impractical.

The detection of ultrasound in photoacoustic imaging has, untilrecently, relied on ultrasonic transducers in contact with thebiological tissue or an ultrasonic coupling agent both of which havemajor drawbacks as described above. Some detection strategies to solvingthe non-contact optical interferometric sensing problems associated withphotoacoustic imaging have been reported.

Optical means of detecting ultrasound and photoacoustic signals havebeen investigated over a number of years; however, to date no techniquehas demonstrated practical non-contact in vivo microscopy in reflectionmode with confocal resolution and optical absorption as the contrastmechanism.

One example of a low-coherence interferometry method for sensingphotoacoustic signals was proposed in U.S. pregrant publication no.2014/0185055 to be combined with an optical coherence tomography (OCT)system, resulting in 30 μm lateral resolution.

Another prior art system is described in U.S. pregrant publication no.2012/0200845 entitled “Biological Tissue Inspection Method and System”,which describes a noncontact photoacoustic imaging system for in vivo orex vivo, non-contact imaging of biological tissue without the need for acoupling agent.

Other systems use a fiber based interferometer with opticalamplification to detect photoacoustic signals and form photoacousticimages of phantoms with acoustic (not optical) resolution. However thesesystems suffer from a poor signal-to-noise ratio, other contact-basedphotoacoustic systems offer significantly improved detectioncapabilities, in vivo imaging was not demonstrated, andoptical-resolution excitation was not demonstrated.

Industrial laser ultrasonics has used interferometry to detect acousticsignatures due to optical excitation of inanimate objects fornon-destructive testing. This approach has been adapted to detectultrasound ex vivo in chicken breast and calf brain specimens, however,optical-resolution focusing of the excitation light was not examined.

Laser Doppler vibrometry has been a powerful non-contact vibrationsensing methodology, however, weak signal-to-noise and poor imagequality have proven to be a limitation when sensing deep-tissue signalsfrom broad-beam photoacoustic excitation.

Similarly, Mach Zehnder interferometry and two-wave mixinginterferometry have been used previously for sensing photoacousticsignals. However many such techniques still require direct contact orfluid coupling; have not offered in vivo studies or optical resolutionfor phantom studies.

BRIEF SUMMARY

According to an aspect, there is provided a photoacoustic remote sensingsystem (PARS) for imaging a subsurface structure in a sample, where thePARS comprises an excitation beam configured to generate ultrasonicsignals in the sample at an excitation location; an interrogation beamincident on the sample at the excitation location, a portion of theinterrogation beam returning from the sample that is indicative of thegenerated ultrasonic signals; an optical system that focuses theexcitation beam at a first focal point and the interrogation beam at asecond focal point, the first and second focal points being below thesurface of the sample; and an interferometer that detects the returningportion of the interrogation beam.

According to another aspect, there is provided a photoacoustic remotesensing system (PARS) for imaging a subsurface structure in a sample,where the PARS comprises an excitation beam configured to generateultrasonic signals in the sample at an excitation location; an opticalsystem that focuses the excitation beam with a focal point that is belowthe surface of the sample; an interrogation beam directed toward anoptical element that is responsive to the ultrasonic signals, theoptical element being interposed between the sample and theinterrogation beam, wherein the generated ultrasonic signals arecharacterized by a returning portion of the interrogation beam; and aninterferometer that detects the returning portion of the interrogationbeam. The optical element may be a Fabry Perot element. Theinterferometer may detect ultrasonic signals to a depth of 7 cm withinthe sample

According to another aspect, there is provided an endoscopic device thatuses a photoacoustic remote sensing system (PARS) for imaging asubsurface structure in a sample, the endoscopic device comprising afiber optic cable having an input end and a detection end; an excitationbeam coupled to the input end of the fibre optic cable, wherein in usethe excitation beam generates ultrasonic signals in the sample at anexcitation location that is adjacent to the detection end of the fiberoptic cable, the fiber optic cable focusing the excitation beam at afirst focal point that is below the surface of the sample; aninterrogation beam coupled to the input end of the fibre optic cable andincident on the excitation location, the fiber optic cable focusing theexcitation beam at a first focal point that is below the surface of thesample, and wherein a portion of the interrogation beam that isindicative of the generated ultrasonic signals is received by thedetection end of the fiber optic cable and travels to the input end; andan interferometer that receives the returning portion of theinterrogation beam at the input end of the fiber optic cable.

According to other aspects, either alone or in combination, asapplicable: the first and second focal points may be within 1 mm of thesurface of the sample; the first and second focal points may be greaterthan 1 μm below the surface of the sample; the focal point may be spacedbelow the surface of the sample at a depth that is greater than a focalzone of the respective at least one of the excitation beam and theinterrogation beam; the excitation beam and the interrogation beam havea lateral separation of less than 1 mm or less than 0.5 mm on thesample; the excitation beam may have a focal point that is laterallywithin the focal zone of the interrogation beam; the interrogation beammay have a focal point that is laterally within the focal zone of theexcitation beam; there may be a processor that calculates an image ofthe sample based on the returning portion of the interrogation beam; atleast one of the first focal point and the second focal point may have afocal diameter of less than 30 μm, 10 μm, or 1 μm; the excitation beammay be scanned through the sample while the interrogation beam isstationary; the interrogation beam may be scanned through the samplewhile the excitation beam is stationary; and each of the interrogationbeam and the excitation beam may be scanned through the sampleconcurrently.

Other aspects will be apparent from the description and claims below.

BRIEF DESCRIPTION OF THE SEVERAL VIEW OF THE DRAWINGS

These and other features will become more apparent from the followingdescription in which reference is made to the appended drawings, thedrawings are for the purpose of illustration only and are not intendedto be in any way limiting, wherein:

FIG. 1 through 4 are block diagrams of optical-resolution photoacousticremote sensing (OR-PARS) microscopy systems

FIG. 5 is a representative drawing of the overlap between the exciterand interrogator beams on a sample.

FIG. 6 is a block diagram of a sensing system involving a Dopplervibrometry configuration.

FIG. 7 is a block diagram of a sensing system using a Fabry-Perotinterferometer.

FIG. 8 is a graph of an example of the response of a Fabry-Perotinterferometer.

FIG. 9 is a block diagram of a Fabry-Perot interferometer.

FIGS. 10A, 10B, and 10C are block diagrams of examples of sensing systemusing a Fabry-Perot interferometer.

FIGS. 11A, 11B, and 11C are block diagrams of examples of sensingsystems in an endoscopy configuration.

FIG. 12 is a block diagram of a sensing system integrated with anotheroptical imaging system.

FIG. 13A is a PARS image of a network of carbon fibres.

FIG. 13B is a graph of the FWHM obtained by fitting an individual carbonfiber signal amplitude to a Gaussian function.

FIG. 13C is a graph of the resolution using a knife edge spreadfunction.

FIG. 13D is a comparison of images obtained by sensing systems inreflection mode and transmission mode.

FIGS. 14A, 14B, 14C, and 14D are in vivo images of CAM-membrane of 5-daychicken embryos.

FIG. 15 is a chart of the measured photoacoustic signals from variousdye concentrations.

FIG. 16 is a graph of the photoacoustic signal vs. excitation energy onthe sample when interrogation power is fixed at 8 mW.

FIG. 17 is a graph of the photoacoustic signal vs. interrogation poweron the sample when excitation energy is fixed at 60 nJ.

FIG. 18A depicts an example of the frequency response of a PARS system.

FIG. 18B depicts an example of the photoacoustic time domain signal ofan individual carbon fiber using a PARS system.

FIGS. 18C and 18D depict an example of the photoacoustic time domainsignal of an individual carbon fiber using a PARS system when theexcitation and interrogation beams are separated by ˜120 and 330 μm,respectively.

FIGS. 19A, 19B, 19C, and 19D depict in vivo PARS images of a mouse ear.

FIG. 20 is a block diagram of a system using a modified version ofpolarization sensitive Michelson interferometry

FIG. 21 is a graph of the measure signal from an unfocused transducer atdifferent lateral distances.

FIG. 22 is a block diagram of a scanning system with two differentinterferometry designs for use with the interrogation beam.

FIGS. 23A, 23B, 23C, and 24 are images of a rat's ear.

DETAILED DESCRIPTION

Photoacoustic imaging is an emerging biomedical imaging modality thatuses laser light to excite tissues. Energy absorbed by chromophores orany other absorber is converted to acoustic waves due to thermo-elasticexpansion. These acoustic signals are detected and reconstructed to formimages with optical absorption contrast. Photoacoustic imaging (PA) hasbeen shown to provide exquisite images of microvessels and is capable ofimaging blood oxygen saturation, gene expression, and contrast agents,among other uses. In most PA and ultrasound imaging systemspiezoelectric transducers have been employed, in which an ultrasoundcoupling medium such as water or ultrasound gel is required. However formany clinical applications such as wound healing, burn diagnostics,surgery, and many endoscopic procedures physical contact, coupling, orimmersion is undesirable or impractical. The system described herein iscapable of in vivo optical-resolution photoacoustic microscopy usingnon-contact optical interferometric sensing without use of anyultrasound medium.

The system described herein, a photoacoustic remote sensing (PARS)microscopy system, is based on the idea of focusing excitation light toa near diffraction-limited spot and detecting photoacoustic signalsusing a confocal interrogation beam co-focused with the excitation spot.While previous approaches used a broad excitation beam with powerfullasers delivering mJ-J of pulse energy over a broad area, the PARSmicroscopy technique described herein uses nJ-scale pulse energiesfocused to near diffraction-limited spots. When focusing into tissue,the surface fluence can be maintained below present ANSI limits forlaser exposure but the ballistically-focused light beneath the tissuecan create fluences transiently far above the ANSI limits (as is done inother microscopy methods). In PARS, this means that very large localfluences ˜J/cm² are created within a micron-scale spot, generating verylarge initial acoustic pressures. For example, at 532-nm excitationwavelength, imaging a capillary with 500 mJ/cm² local fluence wouldresult in an initial pressure on the order of 100 MPa locally. However,because this large pressure is initially localized to a micron-scalespot, by the time the signals are detected by a fluid-coupled detector˜1 cm away, the signals are reduced by 1/r diffractive losses andattenuation to ˜KPa scales. Signals can be orders of magnitude less foracoustic-resolution photoacoustic imaging where unfocused excitationbeams are used, ANSI limits for visible light is 20 mJ/cm², and greaterimaging depths are explored. Large numerical aperture focused acousticdetection is required for optimal signal-to-noise in OR-PAM to ensurethe maximal energy collection. In PARS approach, large optically-focusedphotoacoustic signals are detected as close to the photoacoustic sourceas possible, which is done optically by co-focusing an interrogationbeam with the excitation spot. A long-coherence length interrogationlaser is preferably used with low amplitude and phase noise to read-outthe large local photoacoustic vibrations interferrometrically using anovel architecture designed to optimize received signal intensities.

The high sensitivity and the fine resolution of the proposed systemoffer performance comparable to other in vivo optical resolutionphotoacoustic microscopy systems but in a non-contact reflection modesuitable for many clinical and pre-clinical applications.

Some of the possible options of the optical-resolution photoacousticremote sensing (OR-PARS) microscopy system are depicted in FIG. 1through 4. Variations to the depicted systems will be apparent to thoseskilled in the art. Referring to FIG. 1, a block diagram of PARS system10, and in particular, an optical-resolution photoacoustic remotesensing (OR PARS) microscopy system, is shown. A multi-wavelength fiberexcitation laser 12 is used in multi focus form to generatephotoacoustic signals. Excitation laser 12 preferably operates in thevisible spectrum, although the particular wavelength may be selectedaccording to the requirements of the particular application. Theacoustic signatures are interrogated using a long-coherence length probebeam 16 from a detection laser 14 that is co-focused and co-aligned withthe excitation spots on sample 18. The probe beam 16 passes through abeam splitter 20 that transfers a portion of the signal to a detectionunit 22. Probe beam 16 passes through a polarization control/beamquality unit 24 and a second beam splitter 26 that ties in a referencebeam provider 28. Probe beam 16 then passes through a beam combiner unit30 that also directs excitation beam 32 through a scanning device 34 anda focusing device 36 before reaching sample 18. The reflected beam 38returns along the same path and is analyzed by detection unit 22.

A modified version of polarization sensitive Michelson interferometryhas been employed to remotely record the large local initial pressuresfrom chromophores and without appreciable acoustic loses. The PARSmicroscopy system offers optical lateral resolution down to sub-μm.

Referring to FIG. 2, an example of a PARS experimental setup is shown,with a multi wavelength unit 40 placed in series with the excitationlaser 12 and a lens system 42. Detection laser 14 is placed in serieswith a lens system 42 and a beam splitter 26 that is preferably apolarized beam splitter. As depicted, detection unit 22 is made up of aphotodiode 46, amplifier 48, data acquisition unit 50 and a computer 52.A portion of beam 16 is redirected by beam splitter 26 to a neutraldensity filter 54 after passing through a quarter wave plate 56.Excitation beam 32 and probe beam 16 are combined by beam combiner unit30 and directed to sample 18 by an objective lens 58. Sample 18 may beplaced on a scanning unit 19, allowing for sample 18 to be moved and toallow scanning. FIG. 3 depicts another example of a PARS experimentalsetup using common path interferometry, but without second beam splitter26 or neutral density filter 54. FIG. 3 shows scanning unit 19 placedbetween beam combiner unit 30 and objective lens 58, allowing for thebeam to be moved to allow scanning rather than moving sample 18. FIG. 4depicts a further example of a PARS experimental setup using Michelsoninterferometry, where quarter wave plate 56 has been omitted, and beamsplitter 44 is used to redirect a portion of beam 16 to both neutraldensity filter 54 and detection unit 22.

It will be apparent that other examples may be designed with differentcomponents to achieve similar results. For example, other examples couldinclude all-fiber architectures where circulators replace beamsplitterssimilar to optical-coherence tomography architectures. Otheralternatives may include longer coherence length sources, use ofbalanced photodetectors, interrogation-beam modulation, incorporation ofoptical amplifiers in the return signal path, etc.

The OR-PARS system takes advantage of two focused laser beams on thesample which can simulate a confocal OR-PAM configuration. Since thereare optical components between the objective lens 58 and the sample 18,optical aberrations can be minimized.

Unlike OCT, PARS can take advantage of a high coherence interrogationbeam (HC). In the low coherence interferometry (LC), backscatteringlight is detected from a selected depth (via coherence gating). Howeverin HC method signal from all depth can be detected. Combination of HCdetector with multi-focus excitation improves the SNR.

The OR-PARS takes advantage of optical excitation and detection whichcan help dramatically reduce the footprint of the system. The absence ofa bulky ultrasound transducer makes this all optical system suitable forintegrating with other optical imaging systems. Unlike previousnon-contact photoacoustic imaging systems, the OR-PARS system is capableof in vivo imaging. It relies on much simpler setup and takes advantageof recording the large initial ultrasound pressures without appreciableacoustic loses.

During in vivo imaging experiments, no agent or ultrasound couplingmedium are required. Unlike many other interferometric sensors PARS doesnot require a floating table. No special holder or immobilization isrequired to hold the target during imaging sessions.

PARS can be used to detect ultrasound signals directly. The PARS systemis capable of detecting noncontact measurement of the displacementcaused by ultrasound signals from an ultrasound transducer. In oneexample, a small amount of water was used at the top of the transducerand the transducer was driven by a sine wave from a function generatorat 10 MHz, and produced a noise equivalent pressure of 1 KPa.

Other advantages that are inherent to the structure will be apparent tothose skilled in the art. The embodiments described herein areillustrative and not intended to limit the scope of the claims, whichare to be interpreted in light of the specification as a whole.

A pulse laser is used to generate photoacoustic signals and the acousticsignatures are interrogated using either a long-coherence orshort-coherence length probe beam co-focused with the excitation spots.The PARS system is utilized to remotely record the large local initialpressures from chromophores and without appreciable acoustic loses dueto diffraction, propagation and attenuation.

The excitation beam may be any pulsed or modulated source ofelectromagnetic radiation including lasers or other optical sources. Inone example, a nanosecond-pulsed laser was used. The excitation beam maybe set to any wavelength suitable for taking advantage of optical (orother electromagnetic) absorption of the sample. The source may bemonochromatic or polychromatic.

The receiver beam, or interrogation beam, may be a long-coherence or ashort-coherence length probe beam. In one example discussed above, theprobe beam/receiver beam had a linewidth significantly less than thefrequency of signals detected. Preferably, the interrogation beam has acoherence length selected so that the line width of the laser is lessthan the acoustic signal bandwidth or detection bandwidth

PARS with a long-coherence beam may be integrated with OCT to provide acomplete set of information offered by both photoacoustic and OCTsystems.

PARS with a short or long-coherence beam may be used for either opticalresolution photoacoustic microscopy (OR-PAM) or common photoacousticmicroscopy (PAM).

In one example, both excitation and receiver beam may be combined andscanned. In this way, photoacoustic excitations may be sensed in thesame area as they are generated and where they are the largest. Otherarrangements may also be used, including keeping the receiver beam fixedwhile scanning the excitation beam or vice versa. Galvanometers, MEMSmirrors and stepper/DC motors may be used as a means of scanning theexcitation beam, probe/receiver beam or both.

The configurations shown in FIGS. 5A, 5B, 5C, and 5D may be used toperform PARS imaging. In the depicted embodiments, lines 502, depictedwith a larger radius of curvature, represent excitation beams and lines504, depicted with a smaller radius of curvature, represent receiverbeams. FIG. 5A offers a kind of confocal photoacoustic system where theexcitation beam 502 and probing receive beam 504 are focused on the samespot, which can be on a micron- or sub-micron scale. In FIG. 5B, theoptical resolution can be provided by the receiver beam 504, rather thanthe excitation beam 502. FIG. 5C shows excitation beam 502 and receiverbeam 502 focused on different spots, and takes advantage of ultrasoundtime of flight in order to locate the excitation and receiver beams 502and 504 at different positions. In FIG. 5D, optical resolution isprovided by the excitation beam 502. Preferably, the focus of either orboth of the excitation beam 502 or the detection beam 504 is less than30 μm, less than 10 μm, less than 1 μm, or to the diffraction limit oflight. A tighter focus results in a higher possible resolution and abetter signal to noise ratio in the reflected beam that is detected. Asused herein, the term “focus” is intended to refer to the focal zone ofthe beam, or the point at which the beam spot size is at the tightestsize, and where the diameter of the focal zone is 30% greater than thediameter of the beam spot size. Also preferably, the excitation anddetection beams 502 and 504 are focused on the same position, althoughthere may be some spacing between the respective focuses as shown inFIG. 5C. In FIG. 5C, the beams may be focused at different locations,but preferably within 1 mm, 0.5 mm, 100 μm or within the range of thelargest focus of the beam. In FIGS. 5A, 5B and 5D, the beams may beconfocal, or may overlap within the focus of the beam with the largestfocus. For example, in FIG. 5A, the excitation beam is larger than thedetection beam, and the detection beam is directed at a location withinthe focus of the excitation beam. By moving the detection beam, the areawithin the excitation beam may be imaged. By having confocal beams, bothbeams may be moved to image the sample.

One or both of the beams are preferably focused below the surface of thesample. The beams may be focused, for example, using optics 36 shown inFIG. 1. Generally speaking, the beams may be effectively focused up to 1mm below the surface of the sample. The beams may be focused at least 1μm below the surface, or focused such that focal point of the beam is atleast the distance of focal zone of the beam below the surface of thesample. It will be understood that, while both beams are preferablyfocused below the surface, in some embodiments either the excitationbeam or the interrogation beam may be focused below the surface, withthe other focused on, for example, the surface of the sample. In caseswhere only one beam is focused below the surface of the sample, theseparation between the beams discussed previously will be a lateralseparation, i.e. in the plane of the sample and orthogonal to the depthof the sample.

The excitation beam and sensing/receiver beam can be combined usingdichroic mirrors, prisms, beamsplitters, polarizing beamsplitters etc.They can also be focused using different optical paths.

PARS can be integrated with any interferometry designs such as commonpath interferometer (using specially designed interferometer objectivelenses), Michelson interferometer, Fizeau interferometer, Ramseyinterferometer, Sagnac interferometer, Fabry-Perot interferometer andMach-Zehnder interferometer. The basic principle is that phase (andmaybe amplitude) oscillations in the probing receiver beam can bedetected using interferometry and detected at AC, RF or ultrasonicfrequencies using various detectors. Photoacoustic signals may also bedetected using laser Doppler vibrometry configurations as shown in FIG.6. For clarity, the excitation beam has been omitted. In this example,the detection laser 14 passes through various optical elements,including beam splitters 602 and mirror 604 in order to provide areference optical signal and an optical signal from sample 18 that canbe compared and analyzed by photodetector 606. A Bragg cell 608, orother known device, is used to frequency shift the portion of probe beam16 that is directed to sample 18, as is known in the art.

Another interferometry example is shown in FIG. 7, which uses aFabry-Perot interferometer (FPI) 610 in addition to the PARS systemdescribed above. FPI systems are all-optical detectors that offer highsensitivity and broad bandwidth important for photoacoustic imagingapplications. When used with the PARS system, the FPI may result in animproved sensitivity of the system for deep imaging applications. Asshown, the FPI 610 is reflective to the PARS interrogation beam 16 andvisible to the PARS excitation beam 32 as shown in FIG. 7.

Referring to FIG. 8, an example of the response from FPI 610 is shown,in which FPI resonant peaks are formed in the range of ±30 nm of thePARS interrogation wavelengths. For example if the PARS interrogationwavelength is 1550 nm, the range of the FPI resonant peak would bebetween 1520-1580 nm. The wavelength of the interrogation laser in thePARS system 10 is preferably tunable and will preferably be tuned to thesharpest slope of one the FPI resonant peaks.

Referring to FIG. 9, a detailed view of an example FPI 610 is shown. FPI610 has mirrors 612 and 614 separated by a spacing material 616 andbacked by a backing material 618. The photoacoustic pressure generatedby the excitation beam 32 shown in FIG. 7 causes changes in the opticalthickness and/or physical thickness of the spacing material 616 and/orchanges in the reflectivity of the FPI mirrors. These changes will causea temporary shift of the FPI resonant peaks to a different wavelengthand as a result, changes the amount of interrogation light reflectedback from the FPI to the PARS system. The amount of change in thereflected interrogation light is directly proportional to theultrasound/photoacoustic pressure. The mirrors 612 and 614 of the FPI610 can be fabricated using various techniques, such as glancing angledeposition (GLAD) or other suitable techniques. The spacing material 616may be various materials, such as Parylene C, PDMS, suitable polymers,or other suitable materials. The backing material 618 may be anysuitable material, such as a suitable type of glass or polymer, such asPMMA. The resonant peak of the FPI 610 may be designed to work atvarious wavelengths, such as between 200-2500 nm, depending on therequirements of the application. An Anti-reflection (AR) coating (notshown) can be used on one side or both side of backing material 618 toimprove the light coupling into the FPI 610. It will be understood thatother optical elements that are responsive to an acoustic signal may beused in place of the FPI described above, and that other optical effectsother than those described above may be relied upon to enhance thedetection of the ultrasonic signals in the sample.

As the PARS system 10 generally allows a user to obtain opticalresolution down to sub-μm levels in a non-contact setting, adding a FPI610 can provide acoustic resolution (i.e. greater than 30 μm) detailswith a penetration depth down to 7 cm. Some examples that involve theuse of an FPI 610 are shown in FIGS. 10A, 10B and 10C, based on theorientation of the excitation beam 32 relative to the interrogation beam16. As will be understood, the PARS system 10 may be used intransmission mode, such as in FIG. 10C, using independent generation anddetection units. In the depicted configurations, both excitation andreceiver beams may scan a sample together, one of the beams may bescanning with the other one is fixed, or both beams may be fixed. Inaddition, one or both excitation and receiver beams may be un-focused orloosely focused, or one or both beams may be focused.

As the PARS system takes advantage of an interferometry as explainedherein, in which a reference beam is provided either by an external arm,or in a common mode path. The combination of the interferometry in thePARS system 10 as discussed previously and the FPI 610 can be used toimprove the sensitivity of the photoacoustic imaging system.

In FPI-based photoacoustic imaging systems, the pressure ofphotoacoustic signals change the thickness of the FPI, optically andphysically. These changes cause the shift of FPI resonant peaks andhence change the reflected light from the FPI. In the proposedconfigurations, the FPI 610 is a second interferometer that is addedbetween the PARS system 10 and the sample 18, as shown in FIG. 7. ThePARS system 10 detects surface oscillations in the FPI 610, wherephotoacoustic pressure physically changes the thickness of the FPI 610and in result oscillate the FPI mirrors. The PARS system 10 as thesecond interferometer can pick up the oscillation of the mirrors in theFPI 610. As such, the FPI and PARS detected signal can be combinedtogether to improve the sensitivity of photoacoustic detection. It hasbeen found that, under this system, a detection depth of 7 cm can bereached, although resolution is generally reduced.

In another example, low coherence probe beams can also be considered fordetection of photoacoustic-induced optical phase oscillations but inthis case signals from the sample beam and reference beam will interfereonly if the sample and reference beam paths are equal lengths plus orminus a coherence length (as in optical coherence tomography). It mayalso be beneficial to scan the sample or reference beam path lengths orphase for both range gating and for measuring phase oscillations overthe coherent region of interference.

The reflected light may be collected by photodiodes, avalanchephotodiodes, phototubes, photomultipliers, CMOS cameras, CCD cameras(including EM-CCD, intensified-CCDs, back-thinned and cooled CCDs), etc.The detected light may be amplified by an RF amplifier, lock-inamplifier, trans-impedance amplifier, or other amplifier configuration.Also different methods may be used in order to filter the excitationbeam from the receiver beam before detection. PARS may use opticalamplifiers to amplify detected light prior to interferometry. It mayalso be beneficial to demodulate via Fabry-Perot, very narrow-lineoptical filters, nonlinear and photorefractive crystals and/or spectralhole burning.

An alternative to scanning vibrometry is digital holographic microscopy.Methods of digital holographic microscopy may be used to read outphotoacoustic-induced optical phase oscillations. In this method, apulsed receiver beam may be used so that the sample beam and an angledreference beam interfere on an image sensor (like a CCD or CMOS camera).Using methods of Gabor or Leith-Upatneiks holography not only theamplitude can be recovered, but also the phase of the lightreflected/scattered from the sample over a wide field of view. By gatingwhen the probe beam is pulsed onto the sample relative to the excitationbeam it is possible to stroboscopically reconstruct the photoacousticsignals from each point in the sample as a function of time.Alternatively, it may be possible to time-gate the camera acquisition.The excitation spot may be optically focused or focused over awide-field. When wide-field excitation beams are used,optical-resolution can be achieved by receiving sensing optics and thisresolution is anticipated to depths within a transport mean-free path inturbid media.

PARS can be used in many form factors, such as table top, handheld andendoscopy. Examples of endoscopy PARS are shown in FIGS. 11A, 11B and11C with various arrangements of PARS excitation units 1102, PARSdetection units 1104, fibre optics 1106 such as image-guide fibers, andlenses 1108 that focus the respective beams onto the sample 18. Whenexcitation and detection units 1102 and 1104 are separated, there may bea separate fibre 1110 provided, such as a single mode fiber.

A table top and handheld PARS system may be constructed based onprinciples known in the art. The proposed PARS system takes advantage ofoptical excitation and detection which can help to dramatically reducethe footprint of the system. The footprint of previous systems has beenmuch too large to use the system in all but body surfaces. Forendoscopic applications, the footprint of the ultrasound detector mustbe minimized to make the imaging catheter small and flexible enough tonavigate through small orifices and vessels. The piezoelectric receiversare not ideal candidates for endoscopic applications as there istrade-off between the sensitivity and the size of the receiver. On theother hand for many invasive applications sterilizable or disposablecatheters and a non-contact approach are necessary. The system may alsobe used as PARS endoscopy system with a potential footprint the size ofan optical fiber, as both excitation and PARS beam can be coupled into asingle mode fiber or image guide fiber.

Image-guide fibers (miniaturized fiber bundles with as many as 100,000or more individual micrometer-sized strands in a single optical fiberwith diameters ranging from 200 μm to 2 mm) may be used to transmit bothfocused light spots. The excitation beam may be scanned either at thedistal end or proximal end of the fiber using one of the scanningmethods mentioned before. However, the receiver beam may be scanned orbe fixed. The scanned spot is transmitted via the image-guide fiber 1106to the output end. Therefore, it may be used to directly contact thesample, or re-focused using an attached miniature GRIN lens 1108. In oneexample, C-scan photoacoustic images were obtained from the fiberimage-guides using an external ultrasound transducer to collectphotoacoustic signals. Using an edge-spread and Gaussian function, aresolution of approximately 7 μm was obtained using the image-guidefiber 1106. It is believed that a higher resolution may also be obtainedwith appropriate improvements to the setup and equipment used.

PARS may be used to detect ultrasound signals generated from othersources including ultrasound transducers. The system may also be used asan optical vibrometer. Vibrometers have been used widely for assessingthe operating condition of mechanical properties. Optical vibrometers(OV) offer various advantages over traditional vibration measurementtechniques. The precise metrology of noncontact measurement, highsensitivity and accuracy are the major benefits of optical vibrometers.Most of the common optical vibrometer, including laser Dopplervibrometer (LDV) and Sagnac vibrometer are based on the opticalinterferometry, requiring two coherent light beams. OVs have been usedfor various applications including noncontact measurement of thedisplacement, the acceleration and the velocity of solid surfaces. Thedevice size, cost and noise sensitivity of interferometry arelimitations of the current OVs designs.

In one example, a vibrometry method based on PARS detection was used todetect by noncontact measurement of the displacement caused byultrasound signals from a 10 MHz piezoelectric transducer. The reflectednear-infrared beam from the sample is phase-modulated at the ultrasoundfrequency, and a beat-intensity can be detected. One example had a noiseequivalent pressure of 1 KPa over 10 MHz bandwidth for real-timedetection. This is improved orders of magnitude with lock-in detection.The measurement capability of the system may also be used to measuremotion in high-frequency MEMS actuators and for optical detection ofultrasound.

The PARS system may be combined with other imaging modalities such asfluorescence microscopy, two-photon and confocal fluorescencemicroscopy, Coherent-Anti-Raman-Stokes microscopy, Raman microscopy,Optical coherence tomography, other photoacoustic and ultrasoundsystems, etc. This could permit imaging of the microcirculation, bloodoxygenation parameter imaging, and imaging of other molecularly-specifictargets simultaneously, a potentially important task that is difficultto implement with only fluorescence based microscopy methods. An exampleof a PARS system 10 integrated with another optical imaging system 1202is shown in FIG. 12, where PARS 10 and the other optical imaging system1202 are both connected to the sample 18 by a combiner 1204.

PARS may be used for A, B or C scan images for in vivo, ex vivo orphantom studies.

A multi-wavelength visible laser source may also been implemented togenerate photoacoustic signals for functional or structural imaging.

PARS may be optimized in order to takes advantage of a multi-focusdesign for improving the depth-of-focus of 2D and 3D OR-PARS imaging.The chromatic aberration in the collimating and objective lens pair maybe harnessed to refocus light from a fiber into the object so that eachwavelength is focused at a slightly different depth location. Usingthese wavelengths simultaneously may be used to improve the depth offield and signal to noise ratio (SNR) of PARS images. During PARSimaging, depth scanning by wavelength tuning may be performed.

Polarization analyzers may be used to decompose detected light intorespective polarization states. The light detected in each polarizationstate may provide information about ultrasound-tissue interaction.

APPLICATIONS

It will be understood that the system described herein may be used invarious ways, such as those purposes described in the prior art, andalso may be used in other ways to take advantage of the aspectsdescribed above. A non-exhaustive list of applications is discussedbelow.

The system may be used for imaging angiogenesis for differentpre-clinical tumor models.

The system may also be used for clinical imaging of micro- andmacro-circulation and pigmented cells, which may find use forapplications such as in (1) the eye, potentially augmenting or replacingfluorescein angiography; (2) imaging dermatological lesions includingmelanoma, basal cell carcinoma, hemangioma, psoriasis, eczema,dermatitis, imaging Mohs surgery, imaging to verify tumor marginresections; (3) peripheral vascular disease; (4) diabetic and pressureulcers; (5) burn imaging; (6) plastic surgery and microsurgery; (7)imaging of circulating tumor cells, especially melanoma cells; (8)imaging lymph node angiogenesis; (9) imaging response to photodynamictherapies including those with vascular ablative mechanisms; (10)imaging response to chemotherapeutics including anti-angiogenic drugs;(11) imaging response to radiotherapy.

The system may be useful in estimating oxygen saturation usingmulti-wavelength photoacoustic excitation and PARS or PARS-etalondetection and applications including: (1) estimating venous oxygensaturation where pulse oximetry cannot be used including estimatingcerebrovenous oxygen saturation and central venous oxygen saturation.This could potentially replace catheterization procedures which can berisky, especially in small children and infants.

Oxygen flux and oxygen consumption may also be estimated by usingOR-PARS or PARS-etalon imaging to estimate oxygen saturation, and anauxiliary method to estimate blood flow in vessels flowing into and outof a region of tissue.

The system may also have some gastroenerological applications, such asimaging vascular beds and depth of invasion in Barrett's esophagus andcolorectal cancers. Depth of invasion is key to prognosis and metabolicpotential. Gastroenterological applications may be combined orpiggy-backed off of a clinical endoscope and the miniaturizedPARS/PARS-etalon system may be designed either as a standalone endoscopeor fit within the accessory channel of a clinical endoscope.

The system may have some surgical applications, such as functionalimaging during brain surgery, use for assessment of internal bleedingand cauterization verification, imaging perfusion sufficiency of organsand organ transplants, imaging angiogenesis around islet transplants,imaging of skin-grafts, imaging of tissue scaffolds and biomaterials toevaluate vascularization and immune rejection, imaging to aidmicrosurgery, guidance to avoid cutting critical blood vessels andnerves.

Other examples of applications may include PARS/PARS-etalon imaging ofcontrast agents in clinical or pre-clinical applications; identificationof sentinel lymph nodes; non- or minimally-invasive identification oftumors in lymph nodes; imaging of genetically-encoded reporters such astyrosinase, chromoproteins, fluorescent proteins for pre-clinical orclinical molecular imaging applications; imaging actively or passivelytargeted optically absorbing nanoparticles for molecular imaging; andimaging of blood clots and potentially staging the age of the clots.

PARS MECHANISM

There will now be given a more detailed discussion of the PARSmodulation mechanisms, and comparison with experiment.

The total intensity of light incident on the photodiode is given as theensemble average of the squared magnitude of the sample and referencebeam electric fields (taking constant factors as unity for convenience):

I _(PD) =

|E _(S) +E _(R)|²

The power of light reflected from the surface of the sample then routedto the photodiode is estimated as 0.9²R_(I)I₀ where I₀ is the incidentintensity from the source, and R_(I) is the intensity reflectioncoefficient at the air-sample interface and a factor of 0.9 is includedfor each pass through a 10:90 beamsplitter. Likewise the reference beampower is estimated as I_(R)=

|R_(R)|²

=0.1²η_(VNDF) ²I₀ where η_(VNDF) is the transmissivity of the variableneutral density filter (VNDF).

Possible mechanisms include a pressure-induced refractive-indexmodulation, thermally-induced refractive index modulation, surfaceoscillations, and scatterer position modulation due to confined thermalexpansion. Each mechanism will be discussed below.

Pressure-Induced Refractive-Index Modulation

Refractive index changes due to temperature and pressure rises may inturn affect the scattering of light. We first consider pressure changes.Local initial pressures may be calculated as very large when opticalfocusing and thermal confinement conditions are applied: The initialpressure is given as P₀=ΓΦμ_(a) where Γ is the Gruneissen parameter.Assuming 532 nm light is focused to a micron scale spot size with afocal fluence, Φ, of 500 mJ/cm² and that this light is absorbed byoxygenated blood which has an estimated optical absorption coefficientμ_(a) of 0.0054×43876=236.9 cm⁻¹ (calculated at this wavelength assuminga hemoglobin concentration of 150 g/L and assuming no optical absorptionsaturation) the absorbed energy produces a transient temperature rise ona micro-scale as high as 30K. Using the above parameter estimates, wecalculate an initial pressure as high as 118.5 MPa for unity Grueneisenparameter.

The optical refractive index experiences a perturbation to pressurevariations estimated as

n(r,t)=n ₀(1+ηn ₀ ² P(r,t)/2ρv _(a) ²)

where n₀ is the unperturbed optical refractive index, η is theelasto-optic coefficient (˜0.32 for water), P(r, t) is the pressurefield, ρ is mass density and v_(a) is the speed of sound. Theaccumulated phase shift of light passing through a zone of enhancedpressure can be calculated by Raman Nath diffraction theory and willdepend on the direction of the sound and the direction of the light aswell as the pressure field inhomogeneity. For a light beam incident on aplane pressure wave where both the light and sound beams are parallelthe accumulated phase shift should be zero and are rather maximum whensound fields create effective diffraction gratings orthogonal to thelight propagation. Rather than calculate the phase shifts of transmittedlight we are more interested in the light reflected from a refractiveindex mismatch. With 100 MPa initial pressure a refractive index step ofΔn˜0.019 is predicted in the confined excitation volume, which is a 1.4%change. This results in an optical reflection coefficient of 0.7% whichis very measurable.

This mechanism will contribute to both amplitude and phase variations inthe probe beam.

The electric field back-reflected from the sample and incident on thephotodiode is modelled as having two components, AC and DC terms:

E _(S) =E _(DC,S) +E _(AC,S)

Here, E_(DC,S)=√{square root over (0.9²R_(I)I₀)} is the electric fieldmagnitude of light reflected from the sample surface andE_(AC,S)=√{square root over (0.9 ²I₀t_(I) ²e^(−2μ) ^(eff) ^(d)R_(I,P))}is the electric field amplitude of light reflected from the excitationvolume beneath the surface due to a transient pressure induced opticalindex step. T_(I) is the transmission intensity coefficient at theair-tissue interface. Here both ballistic and scattered photons areaccounted for as reflecting from the index step hence e^(−2μ) ^(eff)^(d) is the effective light attenuation over depth d in a scatteringmedium with effective attenuation coefficient μ_(eff). The factor of 2accounts for a round trip. The fraction of light modulated F_(P) iscalculated as the AC terms from the expansion of I_(PD)=

|E_(S)+E_(R)|²

divided by the DC terms. If the modulated light scattered from theexcitation volume contains multiply scattered photons as we haveassumed, then because phases of the AC component of the sample areeffectively randomized the ensemble average of products of E_(AC,S) withreference beam E_(R) or E_(DC,S) will be zero, leaving the fraction oflight modulated as:

$F_{P} = \frac{\langle{E_{{AC},S}}^{2}\rangle}{{\langle{E_{R}}^{2}\rangle} + {\langle{E_{{DC},S}}^{2}\rangle}}$

Thermally-Induced Refractive Index Modulation

Thermal effects also change refractive indices. Given the volumetricthermal expansion coefficient of water is α_(v)=207×10⁻⁶K⁻¹ at 20° C.and it's refractive index is 1.33, the refractive index would change˜0.01%/° C., hence ˜0.3% with a 30° C. temperature rise, which is smallbut potentially noticeable as a source of scattering modulation.However, this effect would only be applicable locally at the heatingzone. Thermal cooling will occur on a scale ofmicroseconds-to-milliseconds after laser-induced heating.

Surface Oscillations

The above mechanisms point to significant sources of scattering positionor scattering cross-section modulation that could be readily measurablewhen the probe beam is focused to sense the confined excitation volume.However, these large local signals are not the only potential source ofsignal modulation. Acoustic signals propagating to the surface couldalso result in phase modulation of the reflected light. Sound pressuresdecay as 1/r due to diffractive losses so pressure signals can besignificantly weaker away from the confined heating region compared toat their source. For example, assuming an isotropic spherical heatingregion of diameter 5 microns e.g. an isolated red blood cell) the soundpressure level at the boundary of the sphere is estimated as ˜500 MPabut 1 mm away this is predicted as only 2.5/1000*100 MPa. ˜0.25 MPa.Surface pressure modulation can cause surface oscillations and thereflection of the interrogation beam from this oscillating surface canbe a source of detected signal. For a local plane wave at the tissuesurface with peak pressure p the particle velocity is estimated as

$v_{p} = \frac{p}{z}$

where Z is the characteristic acoustic impedance. For a sinusoidalpressure field the particle/surface displacements are thusΔz˜v_(p)/ω_(a) which is a couple of nm for 10 MHz ultrasound with 0.25MPa amplitude. This is a fraction of a wavelength but still asignificant source of phase modulation considering the surfacereflectivity may be high and will return significant amounts of incidentprobe-beam light back to the detector. Notice, however, that thissurface modulation has a 1/ω_(a) dependence hence there is an inherentlow-pass filtering effect that could be rejecting high-frequencycomponents. Nevertheless such signals are readily measurable by oursystem, as evidenced by when we position the probe beam away from theexcitation beam and can still form images when we scan the excitationspot.

If we assume that only surface oscillations are considered, the electricfield phasor from the sample surface is modelled as E_(S)=E_(S0)cos(kzcos(ω_(a)t))˜E_(S0)(J₀(kz)+J₁(kz)e^(jω) ^(a) ^(t)+ . . . ) for acontinuous wave modulation at acoustic angular frequency ω_(a). HereE_(S0)=√{square root over (0.9²R_(I)P₀)}. In this case we do not need toworry about light propagating into the sample, and the fraction of lightmodulated by surface oscillations is estimated as:

$\begin{matrix}{F_{S} = \frac{{2{\langle{{E_{{DC},S}E_{{AC},S}}}\rangle}} + {2{\langle{{E_{R}E_{{AC},S}}}\rangle}}}{{\langle{E_{{DC},S}}^{2}\rangle} + {\langle{E_{{AC},S}}^{2}\rangle} + {\langle{E_{R}}^{2}\rangle} + {2{\langle{{E_{R}E_{{DC},S}}}\rangle}}}} \\{= \frac{{2E_{S\; 0}^{2}{J_{0}({kz})}{J_{1}({kz})}} + {2E_{S\; 0}E_{R}{J_{1}({kz})}}}{{E_{S\; 0}^{2}{J_{0}^{2}({kz})}} + {E_{S\; 0}^{2}{J_{1}^{2}({kz})}} + E_{R}^{2} + {2E_{S\; 0}E_{R}{J_{0}({kz})}}}}\end{matrix}$

Scatterer Position Modulation Due to Confined Thermal Expansion

The volumetric thermal expansion coefficient of water is given asα_(V)=207×10⁻⁶K⁻¹ at 20° C. Assuming a transient temperature rise on amicro-scale as high as 30K as calculated above, a volumetric expansion

$\frac{\Delta \; V}{V} = {{207 \times 10^{- 6}K^{- 1} \times 30K} = {6.21 \times 10^{- 3}}}$

is predicted. For a given temperature rise, the smaller the heatedvolume the larger the expansion. Now if the light is absorbed from a3-micron spot size with a volume modelled as the 1/e penetration depthtimes the cross-sectional illumination area the isotropic particlemotion is modelled as

$\left( {6.21 \times 10^{- 3} \times {\pi (r)}^{2} \times {DOF}_{ex}} \right)^{\frac{1}{3}} = {1\mspace{14mu} \mu \; m}$

which is larger than the wavelength and a very large modulation, where ris the radius of the excitation beam spot size at focus and DOF_(ex) isthe depth of field of excitation beam calculated using Gaussian beamparameters, ˜27 μm.

EXPERIMENTAL RESULTS Experimental Setup

There will now be given an example of experimental methods and setupthat was used to test the principles discussed herein. A modifiedversion of polarization sensitive Michelson interferometry has beenemployed to remotely record the large local initial pressures fromchromophores and without appreciable acoustic loses. The experimentalsetup of the optical-resolution photoacoustic remote sensing (OR-PARS)microscopy system is depicted in FIG. 20. A multi-wavelength visiblelaser source using stimulated Raman scattering (SRS) has beenimplemented to generate photoacoustic signals. Briefly, a 1 ns pulsewidth, frequency doubled ytterbium-doped fiber laser (IPG PhotonicsInc.) with a pulse repetition rate (PRR) of 40 kHz was coupled using afiber launch system (MBT621D/M, Thorlabs Inc.) into a 3 mpolarization-maintaining single-mode fiber (PM-SMF) (HB-450, FibercoreInc., UK) to generate SRS peaks at 543, 560, 590, and 600 nm and pulseenergies up to 500 nJ. The system has been optimized in order to takesadvantage of a multi-focus OR-PAM for improving the depth-of-focus of 2Dand 3D OR-PARS imaging. FIG. 20 shows a tunable pulsed laser source at2202, made up of pulsed laser 12, and having polarization maintainingsingle mode fiber 2204, frequency doubler 2206, collimator lenses 2208,and polarization maintaining nonlinear fibre 2210. Collimator lenses2208 and the second polarization maintaining nonlinear fibre 2210 makeup the SRS peak generation region 2212. After exiting the tunable pulsedlaser source 2202, the beam may encounter an optional band pass filterwheel 2214 before being reflected by mirror 2216. Dual beam combiner2218 passes the beams to galvanometer scanning mirror 2220, throughobjective lens 58, and to sample 18. Galvanometer scanning mirror 2220is controlled by controller and driver 2222, which receives instructionsfrom function generator 2224. Detection laser 14 may be a continuouswave laser, having single mode fiber 2226 followed by collimator lens2208. The beam from detection laser 14 is then passed through polarizedbeam splitter 44. A portion of the split beam passes through lens 42 tophotodiode 2228, through transimpedance amplifier 2230, low-noiseamplifier 2232, band pass filter 2234, to analog to digital convertor2236, providing input for controlling the galvanometer scanning mirror2220, as will be understood by persons skilled in the art. The otherportion of the beam is directed through quarter wave plate 56, through10:90 beam splitter 2238, where the beam is either passed throughvariable neutral density filter 2240 to mirror 2242 to provide referencebeam 2244, or through band pass filter 2246 to combine with the beamfrom the pulsed laser 12 at the dual beam combiner 2218.

The chromatic aberration in the collimating and objective lens pair washarnessed to refocus light from a fiber into the object so that eachwavelength is focused at a slightly different depth location. Usingthese wavelengths simultaneously was previously shown to improve thedepth of field and SNR for structural imaging of microvasculature withOR-PAM. The differences between the multi-focus and single wavelength invivo images will be discussed below.

The output of the PM-SMF was collimated (F280APC-A, Thorlabs Inc.) andcombined using a dichroic beam combiner (DBC) with the receiver arm ofthe system. For the receiver arm a continuous wavelength (CW) C-bandlaser with 100-kHz linewidth (TLK-L1550R, Thorlabs Inc., New Jersey) wasused.

The light at the laser aperture was coupled to a single mode fiber andcollimated. The randomly polarized collimated interrogation beam waspassed through a polarized beam splitter (VBA05-1550, Thorlabs Inc., NewJersey) to be linearly polarized and a λ/4 zero order wave plate(Thorlabs, Inc., New Jersey) to be circularly polarized. The circularlypolarized light then passes through a beam splitter (BS) with 10:90ratio. A variable neutral density filter (NDF) and then a mirror hasbeen placed at the 10% output of the BS in order to provide theoptimized reference power of the interferometry. The beam at the 90%output of the BS has been combined by the excitation arm and thenscanned across the samples via a 2D galvanometer scanning mirror system(GVS012/M, Thorlabs Inc.). The scanning mirrors were driven by atwo-channel function generator. The scanning light was then focusedtightly using an objective lens (M Plan Apo NIR 20X, Mitutoyo, Japan).The reflected light back through the wave-plate creating 90°polarization which then reflects at the polarizing beam-splitter inorder to guide the maximum possible intensity of reflected light to a150 MHz-bandwidth InGaAs photodiode (PDA10CF, Thorlabs Inc., NewJersey). A band pass filter (BPS) has been placed on the detection armto reject the excitation bean. An objective lens (518125, LEICA,Germany) was used in front of the photodiode (not shown in the figure)in order to refocus all possible reflected interrogation light to thesmall photodiode aperture. The output of the photodiode was amplifiedusing an RF amplifier (Olympus 5900PR) with a band pass filter (1 MHz-20MHz) and 26 dB gain and then digitized using a 4-channel PCI digitizer(Gage card) at a sampling rate of 200 MSamples/s. To form images, weproject the maximum amplitude of each A-scan as a pixel in a C-scanen-face image, similar to previous PAM approaches. Since there are nooptical components between the objective lens and the sample (unlikeother reflection mode photoacoustic systems), therefore opticalaberrations can be minimized. Interferometry model for PARS microscopyhas been discussed in the supplementary information section.

FIG. 22 shows two other examples of interferometry designs, having oneof two continuous wave lasers 2404 and 2406. Setup 1, shown at 2410,uses common path interferometry. Continuous wave laser 2404 provides abeam through polarized beam splitter 44, where it is split either towardphotodiode 46, or through quarter wave plate 56 and then to beamcombiner unit 30. Setup 2, shown at 2412, uses Michelson interferometry.Continuous wave laser 2406 provides a beam to 10:90 beam splitter 2238,where it is sent to another photodiode 46, neutral density filter 2408,or to beam combiner unit 30. In both setups shown, pulse laser 12provides an additional beam to polarization maintaining single modefiber 2402, lens system 42, and then to beam combiner unit 30. From beamcombiner unit 30, the beam then passes through objective lens 58 andencounters sample 18.

In a Michelson interferometry, as shown in FIG. 22, a non-polarized beamsplitter 2238 with 10:90 ratio may be used. A variable neutral densityfilter (NDF) 2408 on a 3-axis stage may be placed at the 10% output ofthe beam splitter 2238 in order to provide the optimized referencepower. Like the configuration shown in FIG. 20, the interrogation beammay be combined with the excitation beam and scanned through anobjective lens 58 on the sample 18. For both configurations, anobjective lens (518125, LEICA, Germany, not shown) may be used in frontof the photodiode 46 in order to refocus all possible reflectedinterrogation light to the small photodiode element. The output of thephotodiode 46 may be amplified (Olympus 5900PR) and digitized using a4-channel PCI digitizer (Gage card) at a sampling rate of 200 Samples/s.

Results and Discussion

FIG. 13A shows PARS imaging of carbon fiber networks using ˜1 nJexcitation pulse energy and 6 mW interrogation power. SNR (defined asaverage of signal over the standard deviation of the noise) wasquantified as 45±3 dB. FIG. 13B shows FWHM due to fitting individualcarbon fiber (with ˜6 μm diameter) signal amplitude to a Gaussianfunction. FIG. 13C shows the resolution study using a knife edge spreadfunction. The lateral resolution of the system has been measured as˜2.5±1 μm. FIG. 13D compares the images of reflection mode PARS and atransmission mode OR-PAM using a 10 MHz unfocused transducer (Olympus,V312-SM). The transmission mode OR-PAM setup was not capable of imagingwith 1 nJ pulse energy. FIG. 13D was formed using 50 nJ pulse energy andimages were recorded simultaneously.

In order to validate that the ultrasound signals can be detecteddirectly, the system shown in FIG. 13D was used to analyze a sample 18,which includes a PARS system 10 and an ultrasound transducer 1302 todirectly detect the ultrasound signals generated by the ultrasoundtransducer 1302. The PARS system 10 is also capable of detectingnoncontact measurement of the displacement caused by ultrasound signalsfrom an unfocused piezoelectric transducer (Olympus, A312-10 MHz/0.25″).A small amount of water was used at the top of the transducer 1302 andthe transducer was driven by a sine wave from a function generator at 10MHz. The proposed system has a noise equivalent pressure of ˜1 KPa.

FIGS. 14A, 14B, 14C, and 14D show in vivo images of CAM-membrane of5-day chicken embryos. FIG. 14A shows multi focus PARS images revealingboth capillary beds and bigger blood vessels. In the chicken embryomodel bigger blood vessels usually are located deeper than capillaries.In order to see both deep- and shallow vessels simultaneously themulti-focus design is optimized to extend the depth-of-field to ˜250 μm.FIG. 14B shows a zoomed-in image of both capillary beds and bigger bloodvessels. FIG. 14C shows that PARS is capable of indicating the bleedingarea in the tissue. The bleeding area caused intentionally by using veryhigh pulse energy. FIG. 14D shows PARS images acquired with a singlewavelength (532 nm) rather than multiple wavelengths. With a singlewavelength, depth-of-focus is limited to ˜30 μm, rather than 250 μm withthe multi-focus approach. Hence single-wavelength excitation isbetter-suited for depth-sectioning. This is evident in FIG. 14D, wheretop capillary beds are seen but not deeper large vessels. When we scandeeper with single wavelength excitation we see larger deep vessels butnot superficial capillary beds. A comparison between images in FIGS. 14Aand 14B, with FIG. 14D indicate that multi-focus design helps to improvethe SNR of maximum amplitude C-scan images and improve the depth offield of the system compare to a single wavelength imaging. All theimages shown herein are raw data and no major image processing steps areapplied.

The PARS system is capable of imaging when both beams are scanningtogether, or when the interrogation beam is fixed and excitation beam isscanning. The field of view will be limited in this case as thegenerated photoacoustic signals will experience more attention if theyare located far from the fixed interrogation beam.

FIG. 15 is a chart of the measured photoacoustic signals from variousdye concentrations. FIG. 16 shows the photoacoustic signal (V) vs.energy of the excitation laser (nJ) on the sample when the interrogationpower is fixed at 8 mW. It shows a linear response as expected. FIG. 17shows the photoacoustic signal vs. interrogation power (mW) on thesample when excitation energy is fixed at 600. It is shown that after˜11 mW the photodiode goes to its saturation region.

FIG. 18A depicts the PARS frequency response and FIG. 18B depicts thePARS photoacoustic time domain signal of an individual carbon fiber.Referring to FIGS. 18C and 18D, the charts represent time domainphotoacoustic signal of a single carbon fiber with ˜7 μm diameter whenthe excitation and interrogation beams are separated by ˜120 and 330 μm,respectively. The −3 dB bandwidth was measured as ˜20 MHz by imagingcarbon fiber network as shown in FIG. 13A. The band pass response wasexpected as the RF amplifier was set to a band pass filter (1 MHz-20MHz). The axial resolution of the system is measured ˜75 μm. In FIG. 18b, both excitation and interrogation beam have been co-aligned in the xand y direction and co-focused together in the z direction. Therefore,the photoacoustic signal starts at the time zero. However in FIGS. 18Cand 18D, a time shift has been measured as the excitation andinterrogation beams are separated by ˜120 and 330 μm, respectively. Theresults clearly show the ultrasound time of flight variation by changingthe location of detection spot.

FIGS. 19A, 19B, 19C, and 19D depict in vivo PARS images of a mouse ear,and FIGS. 23A, 23B, 23C, and 24 show in vivo PARS images of a 100 grat's ear. In all in vivo images pulse energy ˜20-80 nJ was used and theinterrogation power was fixed to 6 mW. The selection of excitation andinterrogation lasers will be discussed below.

Unlike OCT, PARS takes advantage of a high coherence interrogation beamin the low coherence interferometry, backscattered light is detectedfrom a selected depth (via coherence gating). However in PARShigh-coherence method signals from all depths can be detected. The depthof images shown in FIGS. 14A, 14B, 14C, and 14D was measured ˜1.5 mm,This may be reduced to ˜1 mm in media with higher turbidity wherefocusing is limited to a transport mean-free path. Aside from depthconsiderations, high-coherence interrogation beams may offer improvedsignal contrast and signal-to-noise compared to low-coherence lasers. Astochastic model is discussed in the supplementary information section.

FIG. 21 shows the measured signal from an unfocused transducer atdifferent lateral distances. The depth of focus of the detection beam onthe sample is measured as ˜50 μm. Therefore, the focus spot diameter ofthe detection beam on the sample is measured as ˜7 μm. We also measuredthe photoacoustic signal generated by different dye concentrations toprove the concept of optical absorption for photoacoustic imaging. It isshown in FIG. 15 that that the sample with higher absorption coefficientprovides bigger photoacoustic signals as expected.

Discussion

In summary the results above showed that: (1) the PARS signal strengthis proportional to optical absorption; (2) the PARS signal strengthincreases linearly with both signal- and reference beam intensities; (3)the PARS signals are largest at the optical focal zone (4) the detectedsignals are indeed photoacoustic signals; (4) signal maximization occurswhen excitation and detection beams are confocal; (5) long-coherence ofthe probe beam is important for signal-to-noise; (6) PARS signaldetection is possible at superficial depths in multiply scatteringtissue; (7) lateral resolution will principally be determined by theexcitation spot size; (8) axial resolution will principally bedetermined by detection system bandwidth; and (9) depth sectioning canbe achieved with high numerical aperture objectives while extendeddepth-of-field can be achieved by harnessing chromatic aberration withmulti-spectral excitation source.

As will be understood, the high sensitivity and the fine resolution ofthe proposed system offer performance comparable to other in vivooptical resolution photoacoustic microscopy systems but with much highersignal to noise ratio and in a non-contact reflection mode suitable formany clinical and pre-clinical applications. In this method amulti-wavelength fiber laser in the visible range has been used in multifocus form to generate photoacoustic signals and the acoustic signatureshave been interrogated using a long-coherence length probe beamco-focused and co-aligned and co-scanned with the excitation spots.

Selecting an excitation laser may involve the following considerations.First, it should be capable of producing suitable conditions of stress-and thermal-confinement that both heat- and stress- build up during thecourse of a laser pulse before the energy can propagate away in the formof thermal diffusion or acoustic propagation. Stress confinement is themost stringent of the two criteria. For example, for 2 μm focusedexcitation spot, laser pulses should be preferably shorter than 2μm/1500 m/s=1.3 ns, which would require a laser with pulse widths of ananosecond or shorter. Second, the repetition rate of the laser willdetermine the imaging frame rate: the faster the repetition rate thehigher. However, the repetition rate is preferably not so high thatsignals from previous pukes overlap in time with subsequent pulses.Given that the imaging depth is not likely to be more than about atransport mean-free path (˜1 mm in tissue) the maximum pulse-repetitionrate is preferably on the order of 1 MHz, Pulse energy should be suchthat ANSI limits are met at the tissue surface, requiring sub-μJ levelsof pulse energy. Finally the wavelength of the excitation source shouldpreferably be tunable for multi-spectral imaging purposes. Theseconsiderations are similar to those for OR-PAM and it is known thatfiber lasers can be a good choice. In one example, a frequency doubledYtterbium-doped fiber laser was used to achieve tenability injectμJ-scale ns-pulses at 532-nm into a length of nonlinear fiber togenerate Stimulated Raman Scattering peaks. A range of wavelengths maybe generated using this technique with enough pulse energy for OR-PAM.other sources are capable of meeting this range of requirements.

The selection of the interrogation laser is also important. As describedabove, the linewidth of the probe laser should be significantly smallerthan the acoustic frequencies to be detected otherwise significant noisepower from the laser source could leach into the passband of the systemand degrade SNR. In one example, a laser was used with a 100 KHzlinewidth, which is significantly smaller than the MHz-level frequenciesto be detected. The laser is preferably tunable in wavelength and power,although wavelength tuning is not critical. A wavelength of 1550 nm maybe used with a 532 nm excitation light because it is spectrallydifferent (important so that optical filters can prevent excitationlight from hitting the detector) and because it is a key band in opticalcommunications where a plethora of components are available. Waterabsorption at this wavelength is higher than desirable: the 1/epenetration depth is a few mm. Other wavelength bands could also beused.

In this patent document, the word “comprising” is used in itsnon-limiting sense to mean that items following the word are included,but items not specifically mentioned are not excluded. A reference to anelement by the indefinite article “a” does not exclude the possibilitythat more than one of the elements is present, unless the contextclearly requires that there be one and only one of the elements.

The scope of the following claims should not be limited by the preferredembodiments set forth in the examples above and in the drawings, butshould be given the broadest interpretation consistent with thedescription as a whole.

1. A photoacoustic remote sensing system (PARS) for imaging a subsurfacestructure in a sample, comprising: an excitation beam configured togenerate ultrasonic signals in the sample at an excitation location; aninterrogation beam incident on the sample at the excitation location, aportion of the interrogation beam returning from the sample that isindicative of the generated ultrasonic signals; an optical system thatfocuses the excitation beam at a first focal point and the interrogationbeam at a second focal point, the first and second focal points beingbelow the surface of the sample; and an interferometer that detects thereturning portion of the interrogation beam. 2-27. (canceled)